Field of the Invention
The invention concerns a method for operating a magnetic resonance imaging system for generating magnetic resonance image data of an object under examination, with which magnetic resonance raw data are acquired. The invention also concerns an actuation sequence for actuating a magnetic resonance imaging system. In addition, the invention concerns a magnetic resonance imaging system designed to implement such a method.
Description of the Prior Art
For the generation of magnetic resonance recordings, the body to be examined is exposed to a relatively high basic magnetic field of for example, 1.5 tesla, 3 tesla, or in newer high magnetic field systems, even 7 tesla and more. Then a suitable antenna device emits a radio-frequency excitation causing the nuclear spins of specific atoms that are excited to resonance by this radio-frequency field in the magnetic field to be flipped by a specific flip angle relative to the magnetic field lines of the basic magnetic field. The radio-frequency signal radiated by the nuclear spins as they relax, called the magnetic resonance signal, is then detected with suitable antenna devices, which can be the same as the transmission antenna device. Following demodulation and digitization, the raw data acquired in such a manner are used in order to reconstruct the desired image data. For spatial encoding of the magnetic resonance signals, respective defined magnetic field gradients are superimposed on the basic magnetic field during the transmission and readout or reception of the radio-frequency signals.
A magnetic resonance recording is typically composed of a number of individual partial measurements in which raw data from different slices of the object under examination are recorded in order subsequently to reconstruct volume image therefrom.
However, in many examinations it is also necessary to perform multiple—i.e. a whole series—of magnetic resonance recordings of the object under examination, wherein a specific measurement parameter is varied. The measurements are used to observe the effect of this measurement parameter on the object under examination in order to draw diagnostic conclusions therefrom later. Here, a series should be understood to mean at least two, but generally more than two, magnetic resonance recordings. Advantageously, in this case, a measurement parameter is varied such that the contrast of a specific type of material excited during the measurements, for example a tissue type of the object under examination or a chemical substance, which is significant for the majority of, or specific, tissue types, such as, for example, water, is affected as greatly as possible by the variation of the measurement parameter. This ensures that the effect of the measurement parameter on the object under examination is particularly clearly visible.
Typical examples of series of magnetic resonance recordings with the variation of a measurement parameter strongly affecting the contrast are diffusion imaging methods (or diffusion weighted imaging (DWI)). Diffusion should be understood to mean the Brownian motion of molecules in a medium. In diffusion imaging, as a rule, multiple images with different diffusion directions and weightings are recorded and combined with one another. The strength of the diffusion weighting is generally defined by the so-called “b-value”. The diffusion images with different diffusion directions and weightings or the combined images derived therefrom can then be used for diagnostic purposes. For example, suitable combinations of the diffusion-weighted images recorded can be used to generate parameter maps providing special diagnostic information, such as, for example, maps reflecting the “apparent diffusion coefficient (ADC)” or “fractional anisotropy (FA)”.
Diffusion imaging is frequently based on echo planar imaging (EPI) due to the short acquisition time of the EPI sequence for each image and its robustness with respect to motion. With diffusion imaging with EPI, even when there is no motion of the patient, which can also play a part, the diffusion-weighted images contain distortion due to local B0 inhomogeneities and residual eddy current fields. The latter are determined by the direction and strength of the diffusion weighting. Such distortion can result in errors in the evaluated diffusion maps.
In the case of diffusion-weighted imaging, additional gradients are inserted into a pulse sequence in order to visualize or measure the diffusion properties of the tissue. These gradients have the result that tissue with rapid diffusion (for example cerebrospinal fluid, CSF) is subject to a greater signal loss than tissue with slow diffusion (for example the grey matter in the brain). The resulting diffusion contrast is becoming increasingly clinically significant and applications now extend way beyond the conventional early identification of ischemic stroke.
A typical pulse sequence for diffusion imaging is the Stejskal-Tanner diffusion sequence, as depicted in FIG. 1. However, when this kind of pulse sequence is used for diffusion-weighted imaging of the liver and heart, signal losses occur which are attributable to the motion of the heart. These signal losses increase with the diffusion weighting and hence result in an overestimation of the apparent diffusion coefficient (ADC). This is a problem when the ADC value is used as a discriminator between benign and malignant lesions.
One possibility for significantly reducing artifacts due to the motion of the heart (and other macroscopic movements), consists in replacing the standard-Stejskal-Tanner diffusion sequence with unipolar gradients by a sequence with velocity-compensated bipolar gradients, as shown in FIG. 2. The main disadvantage of this velocity-compensated sequence is that the bipolar gradients in this sequence have significantly less diffusion sensitivity than the Stejskal-Tanner sequence. Therefore, to achieve the desired diffusion sensitivity, the gradient duration must be extended by at least a factor of 1.6. This results in a prolongation of the echo time of the sequence and hence, due to the inherent T2 decay of the tissue, to a deterioration of the signal/noise ratio (SNR), which can only be compensated by a significantly longer measuring time.
A description of the velocity-compensated diffusion sequence in the prior art and its insensitivity to macroscopic motion can be found in a conference paper given by C. Thomsen, P. Ring and O. Henriksen with the title “In vivo measurement of water self-diffusion by magnetic resonance imaging”, published in “Proceedings of the Seventh Scientific Meeting, Society of Magnetic Resonance in Medicine,” page 890, San Francisco, Calif. (1988).
This sequence was later used in numerous publications in order to reduce artifacts caused by macroscopic motion. A current publication in which the sequence is used in diffusion-weighted imaging of the liver in order to reduce artifacts caused by heart motion, is an article by Masanori Ozaki et al.: “Motion Artifact Reduction of diffusion-Weighted MRI of the Liver: Use of Velocity-Compensated diffusion Gradients Combined With Tetrahedral Gradients”, which appeared the JOURNAL OF MAGNETIC RESONANCE IMAGING, Volume 37, pages 172-178 (2013), DOI 10.1002/jmri.23796. The authors also discuss the aforementioned problem that, compared to a Stejskal-Tanner sequence, for a desired b-value, this sequence requires gradient switching extended by a factor of 1.6 and hence, as a result of the extended echo time, the SNR of the images is reduced. The authors address this by using a direction-independent diffusion preparation with which gradients are switched on all three axes simultaneously. Hence, this achieves reduced TE (and hence improved SNR) compared to a direction-independent preparation sequence with which gradients are only switched on one axis.
Here, reference is made to the fact that stimulated echo preparation (STEAM—from “stimulated echo acquisition mode”), which can achieve a desired diffusion-sensitization with a particularly short echo time, is not velocity-compensated and is hence sensitive to macroscopic motion. Reference is further made to the fact that times of different lengths for diffusion sensitization before and after the RF refocusing pulse are not the exception, but the rule. Symmetrization of the two times (for example due to the non-acquisition of the earlier lines of an EPI readout train) can entail significant problems up to complete signal loss if the echo is shifted in the k-space due to macroscopic motion during diffusion sensitization in the non-acquired range.